Smart Tip LVAD Inlet Cannula

ABSTRACT

Embodiments of the invention provide a left ventricular assist device (LVAD) cannula that includes multiple independent sensors may help decrease the incidence of ventricular collapse and provide automatic speed control. A cannula may include two or more independent sensors. One sensor may measure ventricular pressure, while another may measure ventricular volume and/or ventricular wall location. With this information an automatic control system may be configured to adjust pump speed to minimize the likelihood of ventricular collapse and maximize LVAD flow in response to physiologic demand. Typically the volume sensors are conductance sensors. Further embodiments provide LVADs that are powered by RF energy.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is the national phase under 35 U.S.C. §371 of PCTInternational Application No. PCT/US2013/072611, filed on Dec. 2, 2013,which claims priority to U.S. Provisional Patent App. No. 61/731,879,filed on Nov. 30, 2012, and which is incorporated by reference herein.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with government support under Grant No.HL081119, awarded by the National Institutes of Health. The Governmenthas certain rights in the invention.

BACKGROUND OF THE INVENTION

Field of the Invention

Embodiments of the invention relate to continuous-flow left ventricularassist devices (LVADs) and methods for their use.

Description of the Related Art

A left ventricular assist device or “LVAD” is a mechanical device thatis typically placed in the chest or abdomen of a patient to assist withpumping blood in the patient's body. The LVAD is usually affixed so thatit pumps blood from the left ventricle to the aorta of the patient.Normally an LVAD will include a cable that passes through the skin of apatient to allow control and assessment of the LVAD. The cable alsoallows connection to a controller, power pack, and, typically, a reservepower pack.

Long-term mechanical circulatory support is being used more frequentlyas bridge-to-transplantation and destination therapy for heart failurepatients because of improved safety and reliability of left ventricularassist devices. In addition, mechanically unloading the ventriclereduces wall stress and myocardial oxygen consumption. This can lead toreverse modeling of the myocardium and recovery from heart failure insome instances. Recently, the use of continuous flow assist devices hasbecome common due to their small size and valve-less design.

Unlike pulsatile LVADs, in which pump filling and ejection aredetermined in part by the patient's physiology and inlet cannula suctionpressure is limited by atmospheric pressure, continuous flow LVADsproduce a flow-dependent differential pressure as a function of pumpspeed as described by the characteristic pressure versus flow (H-Q)curve. If the speed is too slow, the patient may not receive an optimalamount of blood flow and their activities may still be limited by theirheart failure. If the pump speed is too fast, the pump can empty theventricle, pulling the ventricular wall towards the pump inlet andsubsequently limiting flow. This phenomenon, referred to as a suctionevent, can cause myocardial damage and dangerous ventriculararrhythmias.

Infrequent but occasional aortic valve opening is often used as aguideline to set pump speed, indicating the ventricle has adequateresidual volume at end-systole to prevent suction and is unloaded tosome extent. However, the aortic valve opening is only measured in theclinic setting and is subject to changes in left ventricular (LV)contractility, heart rate, arterial pressure, and blood volume relatedto normal daily activities (e.g. sleep, exercise, positional changes,etc.). An automatic control algorithm is desirable to adjust pump speedin response to hemodynamic changes in order to provide sufficientsupport but reduce the risk of suction-induced arrhythmogenesis.

Of the many methods that have been devised to control continuous flowLVADs, most rely on an estimate of instantaneous pump flow based on themotor equation and power dissipation. However, accuracy is affected bythe blood viscosity, model non-linearities, and noise in the powermeasurement. In addition, pump flow alone cannot determine theventricular workload, and therefore, cannot be used to optimize pumpspeed to unload the ventricle. A control algorithm based on directmeasurement of left ventricular volume and/or left ventricular pressurewould be advantageous in preventing suction events and setting anoptimal operating point that reflects ventricular loading. Onceestablished, this control system can be used with existing pumps usedclinically and allow adaptive flow control that can adjust withphysiologic demand (i.e. changes in ventricular load). As LVAD patientsleave the hospital and return to their daily activities, the controlsystem will be able to adapt to the patient's circulatory needs andprevent adverse suction events.

BRIEF SUMMARY OF THE INVENTION

We have found that an LVAD that includes independent sensors formeasuring at least one of ventricular pressure and volume may helpdecrease the incidence of ventricular collapse. We refer to this as a“smart” inlet cannula with integrated sensors for measuring pressureand/or volume of the left ventricle (“LV”). An automatic control systemcan use the volume and pressure information to adjust pump speed tominimize the likelihood of collapse. In addition, the control system canuse the sensor information to assess circulatory needs (e.g. duringexercise), and adjust pump speed accordingly. The sensors areconductance electrodes for measuring volume and a pressure sensor.Further embodiments provide LVADs that are powered by RF energy.

DETAILED DESCRIPTION OF THE FIGURES

FIG. 1 shows an engineering design of an embodiment of the invention.

FIG. 2 shows an inlet cannula with integrated pressure and volumesensors. Data may be transmitted wirelessly to the VAD controller, whichmay be external or internal.

FIG. 3 shows a pressure sensor.

FIG. 4 details electrodes of a conventional conductance catheter.

FIG. 5 shows a block diagram of the smart tip internal electronics.

FIG. 6 shows a block diagram of the external transceiver.

FIG. 7 shows left ventricular pressure-volume relationship as a functionof LVAD pump speed.

FIG. 8 shows signals recorded from the smart tip pressure sensorcompared to a reference Millar transducer during an acute ovine study.

FIG. 9 shows signals recorded from the smart tip volume sensor comparedto the LV short axis dimension measured using sonomicrometry during anacute ovine study.

FIG. 10 shows signals recorded from the smart tip conductance sensorcompared with the LV short axis dimension as pump speed increasedleading to suction during an acute ovine study.

FIG. 11 shows an example of suction detection based on inlet pressure inan acute calf study.

FIG. 12 shows and example of LV unloading control based on inletpressure in an acute ovine study.

DETAILED DESCRIPTION OF THE INVENTION

We have found that a left-ventricular assist device (“LVAD”) thatincludes multiple independent sensors may help decrease the incidence ofventricular collapse during LVAD use. Embodiments of the invention mayinclude conductance electrodes and/or pressure sensors in an LVADcannula tip for determination of ventricular pressure, ventricularvolume, and ventricular wall location.

The sensors are conductance and/or pressure sensors. The use ofconductance electrode technology may be better understood with referenceto the figures. FIG. 1 shows an embodiment of the invention in which awall is shown as transparent. This transparency is for the convenienceof the viewer in FIG. 1, and is not a requirement of embodiments of theinvention. FIG. 1 shows an LVAD cannula tip including an outer wall 1surrounded by a plurality of concentric electrodes 3 in communicationwith electronics 17 and a transceiver coil 5. In a preferred embodimentthere are four concentric platinum electrodes, though those of skill inthe art will, with the benefit of this disclosure, recognize thatadditional numbers and types of electrodes are possible. For example,electrodes may be made from titanium with a thin layer of platinum blackon the blood contacting surface.

The transceiver coil is located in a cavity 7 defined by an inner wall 9and the outer wall. The inner wall also defines an opening 11 for theLVAD inlet cannula. A pressure transducer 12 is mounted between theinner and outer walls, with the sensing surface in communication with asmall sensing chamber filled with silicone gel. One side of this chamberis formed by the inner wall where the wall thickness is very thin, suchthat blood pressure on the inner wall is transmitted to the sensingchamber and thereby to the pressure sensor. The cannula tip is connectedto a cannula (a flexible tube that carries blood) which then connects tothe LVAD inlet port. In another embodiment the cannula tip may beincorporated directly in to the LVAD inlet port, such that a cannula isnot used.

In a preferred embodiment, the pressure sensor communicates withelectronics that provide functions such as signal conditioning,excitation, and analog-to-digital conversion or other processing. Atransceiver coil enables wireless transmission of data to an externalreceiver and LVAD controller, and also functions as a receiver forwireless power coupled inductively from a power transmitting coil placedoutside the body. Separate coils may also be used for receiving power,receiving data, and transmitting data. Electronics may also providefunctions related to power reception and conditioning. Electronics mayalso be located in a separate location, or a portion of the electronicsmay be located in the tip, and power and data may be transmitted overcables.

In a typical embodiment of the device, four electrodes are used forconductance measurements. In this arrangement a time-varying currentsource is applied through the outer electrode pair, and voltage ismeasured across the inner pair of electrodes. Separating the measurementelectrodes from the source electrodes eliminates measurement errors dueto the polarization potential at the electrode-blood interface.

The conductance that is measured when the LVAD cannula is used will be acombination of the conductance of the blood (the desired blood volumemeasurement) and the surrounding tissue. The conductance of thesurrounding tissue may be compensated for by a dual frequency method orby the methods of Baan et at or Wei et al. Determination of conductanceis also reported in more detail herein.

LVADs of embodiments of the invention may include or be connected to anautomatic control system. Such a control system may use data obtainedthrough the conductance electrodes and/or pressure sensor to adjust pumpspeed to minimize the likelihood of ventricular collapse and to maximizepump flow in response to circulatory demand. If the measured volume isbelow a predetermined threshold, pump speed will be reduced. Likewise,if the measured volume is above a predetermined threshold, pump speedwill be increased.

In further embodiments of the invention the LVAD is powered by RFenergy. This is accomplished, for example, by inductive coupling betweentwo coils of wire. The internal coil is implanted under the skin andconnects to power conditioning circuitry contained in an implantablemetal container, similar to a pacemaker. The conditioned power is usedto drive the implantable blood pump, and may also recharge implantablebatteries to be used whenever power transfer is absent. The externalcoil is placed against the skin in proximity to (usually within 1 inchof) the internal coil. The external coil is connected to a poweroscillator and controller, which produces current in the external coil,and thereby induces current in the internal coil. In most cases thecoils are tuned, via capacitors connected in series or parallel with thecoils, to resonate at a given frequency which improves power transferefficiency.

The following information is provided to give a more robust descriptionof various embodiments of the invention, which are described in afeature-by-feature basis. Those skilled in the art will appreciate thatthese embodiments are exemplary, that they suggest many possiblecombinations, and that the invention is defined by the claims.

Pressure Sensor Design

The requirements for direct pressure measurements are challenging. Thesensing element must be resistant to thrombus formation, which requiresnon-thrombogenic materials and the absence of crevices, steps, or othersurface features. In one embodiment we measure LV pressure relative toatmospheric pressure, as in conventional catheterization P-V loops. Forlong term VAD implantation, which is now approaching five years in somecases, offset drift should be minimal, although some compensation fordrift may be accomplished in a final control algorithm. The frequencyresponse, for control purposes, should be approximately 0-100 Hz.

Pressure Sensor

We have developed a pressure sensor to measure inlet cannula pressure.We have used a pressure signal in a VAD speed controller for avoidingleft ventricular suction, both in vivo and in vitro. A strain gagebonded to an area of the pump inlet port that is machined to between0.001 and 0.005 in thick, for example, provides a signal that isproportional to pressure in the tubing lumen. Finite element modeling(FEA) was used to develop the sensor. This approach provides adequatesignal-to-noise ratio, but long term offset drift was suboptimal. Thiswas due to the extremely low strain and the use of individually bondedstrain gauge elements directly to a titanium diaphragm. Strain producedby the mismatch in the temperature coefficients of expansion of thesilicon, titanium, and adhesive, as well as creep and curing stresses inthe adhesive, produced error signals on the same order of magnitude asthe pressure-produced strain of interest. The sensor gain was stable,however, such that control algorithms based on the pulsatility (max-min)of LV pressure would be possible.

To reduce offset drift described above, we have developed a pressuresensor coupled to a thin-walled section in the inlet cannula lumen.Silicone dielectric gel or silicone oil is used as a couplant betweenthe sensor and the cannula, as shown in FIG. 3. FIG. 3 shows a pressuresensing region 23, silicone gel or oil 25, and a thin wall flexingportion 27 of an inner lumen.

A number of candidate pressure sensors, with similar performancespecifications but different mounting options may be suitable for use inembodiments as reported herein. For example, the MS58XX series sensors(Measurement Specialties, Inc., Hampton, Va.) consist of a siliconmicromachined pressure sensor die mounted on a 6.2×6.4 mm ceramiccarrier or a 6.1×6.3 mm PCB protected by a metal or plastic cap. Thesensor element consists of a micro-machined silicon membrane with Pyrexglass waferbonded under vacuum to the back side for reference pressure.Ion-implanted resistors make use of the piezoresistive effect to sensepressure applied to the membrane. The sensor is mounted using a processallowing offset stability, making the device suitable for direct PCBassembly. An advantage of this approach is that the offset driftassociated with strain gauge bonding is controlled and minimized in thesensor fabrication process. Final assembly of the cannula does notrequire direct bonding of individual strain gauge elements or die.

Assuming an LV pressure range of −50 to 150 mmHg (gauge), and barometricpressure ranging from normal 1 atm at sea level to the typical aircraftcabin pressure equivalent of 9000 ft elevation, a typical absolutepressure range is given in Table 1.

TABLE 1 Pressure sensor operating range Absolute Absolute Pressure @Gauge Pressure @ 9000 ft elevation Pressure sea level (1 atm) (0.75 atm)Minimum −50 mmHg 710 mmHg 520 mmHg (0.934 atm, 94.7 kPa) (0.684 atm,69.3 kPa) Maximum 150 mmHg 910 mmHg 720 mmHg (1.197 atm, 121 kPa) (0.947atm, 96.0 kPa)

Based on specifications of the MS58XX series with maximum pressures inthe range of 1 to 1.3 bar absolute, long term offset drift is specifiedas less than 1 mmHg per year.

Every sensor is individually factory calibrated at two temperatures andtwo pressures. As a result, 6 coefficients necessary to compensate forprocess variations and temperature variations are calculated and storedin the 128-bit PROM of each module. These bits (partitioned into 6coefficients) are read by the microcontroller software and used toprovide compensated pressure and temperature values:

C1: Pressure sensitivity

C2: Pressure offset

C3: Temperature coefficient of pressure sensitivity

C4: Temperature coefficient of pressure offset

C5: Reference Temperature

C6: Temperature coefficient of the temperature sensitivity

Membrane strain estimates

The pressure sensor is not in direct contact with blood, but is coupled(via incompressible silicone gel or silicone oil) to a thin-walledsection of the cannula lumen, which acts as an isolation diaphragm.Because of the small size and high sensitivity of the MEMS pressuresensing area, the required deflection of the isolation diaphragm isextremely small.

One embodiment of the pressure sensor using the MS5803 sensor and a0.005 in diaphragm was tested in an acute ovine study. Signals from thesmart tip pressure sensor are shown in FIG. 8 compared to a referencesensor implanted in the ventricular cavity (Millar Instruments, Inc.)

Conductance Electrodes

Volume conductance catheters are commonly used in cardiac studies inanimals, and occasionally in humans, to provide continuous measures ofventricular volume. Concurrent LV pressure measurements yield P-V loops.As shown in FIG. 4, a pair of constant current conductance electrodes 29and 31 having excitation and sensing ends create an electric fieldwithin a ventricle 33 along a recording segment 35. The electricpotential measured at the sensing electrodes is roughly the product ofthe current density and the total conductance between the electrodes (aswell as the uniformity of the electric field).

The original equation for volume by Baan, et al. is

${Volume} = {\frac{1}{\alpha}\rho \; {L^{2}\left( {G_{meas} - G_{p}} \right)}}$

ρ—blood resistivity

L—length between voltage sensing electrodes

α—a calibration factor

G_(meas)—conductance measured

G_(p)—parallel conductance of muscle

Gp is measured by using a hypertonic saline bolus method, and thecalibration factor a is determined by measuring stroke volume (SV),usually with a flow probe on the aorta in animal studies

An improved method developed by Wei et al. uses the complex conductance(i.e. admittance) to separate muscle from blood, since myocardiumexhibits both resistance and capacitance, while blood is purelyresistive. This improved method reduces errors due to position of thecatheter in the ventricle. Wei et al., developed a non-linear equationfor volume where y is a field form factor; this approach demonstratedimproved accuracy even with radial misalignment of the catheter. Themodified equation for volume is

${Volume} = {\frac{1}{1 - {G_{b}/\gamma}}\rho \; {L^{2}\left( G_{b} \right)}}$G_(b) = Ysin (θ) − G_(m)$\; {G_{m} = {C_{m}\frac{\sigma_{m}}{ɛ_{m}}}}$$C_{m} = \frac{{Y}{\sin (\theta)}}{2\; \pi \; f}$

The constants σ_(m) and ε_(m) are the muscle conductivity andpermittivity, respectively. |Y| is the magnitude of the total measuredconductance at an excitation frequency f and phase angle θ. In VADpatients, echocardiography, which is routinely performed in VADpatients, will be used to measure end-diastolic and end-systolicadmittance in order to compute the field form factor γ, as per Wei etal.

The proposed electrode configuration, in the apical region only, willonly be sensitive to blood volume near the cannula tip. However, thetotal LV volume is expected to be closely related to apical volume, fora given patient, and γ will therefore include correction for thedifference between measured apical volume and total volume. Furthermore,for the suction avoidance control, the volume near the cannula tip ismore relevant than total volume.

An advantage in this application is the large diameter of the inletcannula, compared to the small diameter of conductance catheters; theincreased electrode surface area will provide reduced electrodeimpedance.

The calculation of C_(m) may also be used to detect misalignment of theinlet cannula; by providing a measure of muscle proximity to the cannulatip.

Electrode Design

A separate pair of electrodes is usually used for constant currentexcitation, and one or more pairs of electrodes are used to measure theconductance-dependent voltage. The reason for separate excitationelectrodes is that the current density associated with excitation causesa polarization potential, which adds to the desired voltage measurementdue to volume conductance. By separating excitation from sensing, thesignal-to-noise ratio and signal stability are improved, at the expenseof an additional pair of excitation electrodes.

As shown in the embodiment of FIG. 2, the cannula may be built withtetrapolar electrodes, using four platinum or platinum-iridium ringelectrodes, installed into grooves on a PEEK cannula body. Signalsrecorded from the smart tip conductance sensor during an acute ovinestudy are shown in FIGS. 9 and 10. The conductance measurement is showncompared to the ventricular short axis dimension obtained usingsonomicrometry. Conductance increases/decreases as the LV axis dimensionincreases/decreases. FIG. 2 shows an outer wall 1, inner wall 9,pressure sensor 19, conductance electrode 21, and electronics suite 17.A RF coil 22 is also shown. Blood flow direction is also shown in FIG.2.

Further embodiments may eliminate the two inner electrodes by using theouter electrodes as combined excitation and sensing electrodes. Theeffect of the polarization potential can be reduced by using higherexcitation frequency, and lower current density.

Historicallly, there has been an increase in excitation frequency fromtens of kHz to 100 kHz. In a typical smart cannula application accordingto an embodiment of the invention, the electrode surface area issignificantly higher than in conventional conductance catheters. Oneembodiment uses a user defined excitation frequency that can varybetween 1 and 100 kHz.

Typical values for excitation current in conventional conductanceelectrodes are in the microamp range. Low current density is desired tominimize the effect of the polarization potential, due to limited ionmobility at the electrode-tissue interface. In the current embodiment,the larger surface area electrodes allow increased excitation currentsto be used (e.g. 0.5 to 5 mA) used while maintaining low currentdensity. The increased excitation current improves the signal to noiseratio in the recorded signal.

The excitation frequency, magnitude, and bipolar/tetrapolar arrangementmay be selectable through telemetry, and software controlled in thecannula. In various embodiments it will be possible to rapidly switchthese modes, even during a cardiac cycle, to compare the effect ofbipolar vs tetrapolar, excitation frequency, and excitation currentmagnitude.

Electronics, Power, Telemetry

One option to access the pressure sensor and conductance catheter wouldbe by direct wiring, ie. a miniature cable from the cannula to the VADcontroller. However, implantable electronics, with wireless telemetry ofthe pressure and volume signals, offers a number of advantages in termsof development and staged integration with VAD controllers. Advantagesmay include, for example improved ease of integration with existingcontrollers without necessitating an additional percutaneous line.

System Architecture

FIG. 5 shows an implanted portion of one embodiment of the smart cannulaelectronics. Because of the current level of miniaturization in surfacemount packaging, driven by the market for portable electronics (cellphones, iPods, laptops, etc.) embodiments may be created without the useof custom ICs. One may use custom and semi-custom options, such asASICs, which can integrate most of the functions on a single chip.

Power

In one embodiment, an LVAD of an embodiment of the invention is poweredas follows:

-   -   Class E transceiver    -   1-2 MHz    -   parallel tuned secondary    -   Highest load requirements ˜10 mA, 3.3.V, 330 mW.

Outgoing Telemetry

In one embodiment, useful outgoing telemetry transmitted by the LVADincludes the following:

Baseband data: serial, 3 channels (Pressure, Conductance mag,Conductance phase) 100 samples/sec each=300 samples/sec

12 bits per sample=3600 bits/sec

Outgoing-impedance modulation at 2 FSK frequencies centered at a minimumof 10 times bit rate=36000 Hz; e.g. FSK frequencies of 40 kHz and 60 kHz

Detection by tracking frequency changes in Class E driver. Need FSKbecause frequency will shift with coupling etc. Class E controller needsto be fast enough to track 40-60 kHz, or track slowly and detect thephase error, as long as the error is small and does not impactefficiency or EMI.

Ingoing Telemetry

In embodiments of the invention, inbound telemetry received by an LVADmay include the following

Baseband data: very slow (e.g. 10 bits/sec), used to selectbipolar/tetrapolar, and excitation magnitude and frequency.

ASK may be obtained by modulating the Class E drive magnitude at ASKfrequency of 0 Hz and 100 Hz; detected at implant by simple peakdetector, demodulate in implant software using timer.

External Transceiver

External transceiver functions are shown in FIG. 6.

Sensor Scaling and Correction

The pressure sensor may include coefficients for one or both of pressureand temperature compensation, which will be calculated by an internalmicroprocessor.

Barometric pressure will be measured externally, using the same sensorfamily as used internally, and the barometric pressure will besubtracted from the absolute pressure measure from the smart cannula.

LVAD Speed Calculation

Suction Avoidance

Using the inlet conductance catheter an algorithm has been developed tocontrol the pump speed of continuous flow LVADs. There are two aims ofthe control system: (1) immediate speed reduction following detection ofsuction, and (2) beat-to-beat control of pump speed based on enddiastolic volume (EDV).

LV volume will be continuously monitored to detection suction. Suctioncan occur if the pump speed is too high or there is insufficient bloodflow return to the LV. A threshold value will be set empirically, and adrop in LV volume below the threshold level will trigger an immediatereduction in pump speed until the suction event is resolved.

Suction detection using the smart tip conductance electrodes during andacute ovine study is shown in FIG. 10. Conductance recordings over 1cardiac cycle are shown as speed increased from 7000 to 11700 rpm. At11400 rpm suction occurred as can be seen as a reduction in themagnitude and pulsatility of the conductance signal. The LV short axismeasured using sonomicrometry was used as a reference for the study.

A suction detection algorithm has been developed using inlet pressure.Suction is readily detectable as a highly negative, short durationsignal (i.e. large -dP/dt) from the pressure sensor. Smaller amplitudenegative pressure transients occur at the beats leading up to completesuction. An example of suction detection in an acute calf study is shownin FIG. 11. Pump speed was increased until suction was detected by dP/dtexceeding a pre-determined threshold. Following suction detection, pumpspeed was automatically reduced and the suction event was eliminated.After resolution of suction, pump speed was gradually increased to theset point. The control system was able to detect suction events andreduce pump speed immediately to resolve the suction event.

LV Unloading Control

In the absence of suction events the control system will adjust pumpspeed to optimize pump flow and reduce LV pressure and volume loading.Unloading will be set by measuring the end diastolic volume (EDV) on abeat-to-beat basis and adjusting the pump speed accordingly as depictedin FIG. 7. An increase in EDV with activity, etc. will cause anappropriate increase in pump speed to provide more unloading andcirculatory support. A decrease in EDV may occur during rest or sleep,and the pump will reduce speed accordingly to prevent suction. It isexpected that the initial EDV set point, and a EDV set point range, willbe determined for a given patient based on echocardiographic assessmentof LV volume and the frequency of aortic valve opening. Weaker heartswill tend to have lower EDV set points. As compared to current fixedspeed pump control system, the proposed control system will provide pumpflow that adapts to physiologic demand.

In addition to using EDV to adjust pump speed, we have developed acontrol system to adjust speed based on either LV end-diastolic pressure(EDP) or peak-to-peak pressure (Pppk) to assess cardiac preload. Pppk isdefined as the difference between the end-systolic and end-diastolicpressures over each cardiac cycle. Using this relative pressuremeasurement mediates the possibility of baseline drift of the pressuresensor signal. The control system was tested in an acute sheep study,and the results are shown in FIG. 12. Initially, the control system setthe pump speed to 8000 rpm. Partial occlusion of the IVC reduced LV EDPand Pppk, and the control system responded by decreasing pump support.The system stabilized at a reduced pump speed and LV EDP and Pppkreturned to baseline. The occlusion was then released causing anincrease in LV EDP and Pppk, and the control system responded byincreasing pump support.

Long Term Management of LV Conditioning—Assessing Contractility fromDirect Measurement of Pressure-Volume Relationship

Evaluation of ventricular function is essential to determine myocardialrecovery and weaning from pump support. End-systolic elastance (E_(es))and the end-diastolic pressure volume relationship (EDPVR) are the goldstandards for assessing native heart systolic and diastolic function,respectively (Suga and Sagawa 1974). Indirect assessment techniques ofsystolic function have been developed due to the inability to measurethe pressure-volume relationship directly (Endo, Araki et al. 2001;Kikugawa 2001; Nakata, Shiono et al. 2001; Naiyanetr, Moscato et al.2009). These methods rely on current or flow measurements and theinverse calculation of ventricular pressure and volume that requiresmultiple assumptions. In addition, these techniques are not able toassess diastolic function. With the proposed cannula the pressure-volumerelationship of the native ventricle will be measured directly, whichwill enable the use of E_(es) and EDPVR to assess ventricular function.

Traditional measurement of E_(es) and EDPVR require venal occlusion togenerate serial pressure-volume loops. To enable non-invasivemeasurements, Sensaki et al. and Shishido et al. developed single-beatmethods to assess E_(es) without the need for serial pressure-volumedata. The methods use a normalized elastance function that is fit to thesingle pressure-volume loop during the isovolumetric contraction andejection phases. Sensaki et al. reported the single-beat method as onethat provides a reliable estimate of contractility in humans that isminimally affected by loading conditions. The method has been showncapable of estimating native contractility even during assistance withrotary blood pumps. Recently, a similar single-beat method has beenreported to estimate EDPVR (Klotz, Hay et al. 2006). Single-beatestimation of E_(es) and EDPVR has been proven in heart failure patientsto be an effective clinical tool.

Using a single-beat method, E_(es) and EDPVR are obtained periodicallyfrom the pressure-volume loop data. A baseline systolic and diastolicfunction level are determined for each patient at initial pump support,and improvement in systolic and diastolic function are determinedrelative to the baseline level. To minimize variability due to pumpsupport, the pump speed will be fixed during baseline evaluation and allsubsequent functional assessments. Absolute functional levels are notrequired to assess functional recovery, and therefore, errors in thesingle-beat method will be mitigated. By providing a real-time estimateof myocardial functional improvement, clinicians will be able toevaluate and optimize unloading strategies in order to enhance recoveryand device weaning.

Physical Integration of the Smart Cannula into Existing or Future LVADs

To integrate the smart cannula with existing LVADs the physicaldimensions of the device would have to change but the basic design wouldnot be affected. Physical integration of one embodiment of a smartcannula into existing or future LVADs is straightforward.

Any documents referenced above are incorporated by reference herein.Their inclusion is not an admission that they are material or that theyare otherwise prior art for any purpose.

1. A tip for attachment to a left ventricular assist device, comprising:an inner wall and an outer wall having a distance between them defininga cavity, wherein: said inner wall further defines a passage formovement of blood there through, said inner wall further defines anopening for an inlet cannula of the left ventricular assist device, saidinner wall further defines a sensing chamber at a portion of said innerwall that has a reduced thickness; at least one conductance sensordisposed about the outer wall; and at least one pressure sensor disposedin the cavity, the pressure sensor in communication with the sensingchamber.
 2. The tip of claim 1, wherein said at least one conductancesensor comprises a series of electrodes.
 3. The tip of claim 2, whereinthe series of electrodes have a shape selected from the group consistingof ring-shaped and patch-shaped.
 4. The tip of claim 1, wherein said tipis in communication with a radio frequency power supply for operation ofa left ventricular assist device.
 5. The tip of claim 1, wherein saidtip is in communication with a control system.
 6. A tip for attachmentto a left ventricular assist device, comprising: an inner wall and anouter wall having a distance between them defining a cavity, whereinsaid inner wall further defines a passage for movement of blood therethrough; and at least one conductance sensor disposed about the outerwall.
 7. The tip of claim 6, wherein said at least one conductancesensor comprises a series of electrodes.
 8. The tip of claim 7, whereinthe series of electrodes have a shape selected from the group consistingof ring-shaped and patch-shaped.
 9. The tip of claim 6, wherein said tipis in communication with a radio frequency power supply for operation ofa left ventricular assist device.
 10. A tip for attachment to a leftventricular assist device, comprising: an inner wall and an outer wallhaving a distance between them defining a cavity, wherein said innerwall further defines a passage for movement of blood there through,wherein said inner wall further defines an opening for an inlet cannulaof the left ventricular assist device, wherein a sensing chamber isformed at a thinned portion of said inner wall; and at least onepressure sensor disposed in the cavity and in communication with thesensing chamber.
 11. The tip of claim 10, wherein said at least oneconductance sensor comprises a series of electrodes.
 12. The tip ofclaim 11, wherein the series of electrodes have a shape selected fromthe group consisting of ring-shaped and patch-shaped.
 13. The tip ofclaim 10, wherein said tip is in communication with a radio frequencypower supply for operation of a left ventricular assist device.
 14. Thetip of claim 10, wherein said tip is in communication with a controlsystem.
 15. A left ventricular assist device comprising the tip ofclaim
 1. 16. The left ventricular assist device of claim 15, furthercomprising a control system and a radio frequency power supply incommunication with said tip.
 17. A method for adjusting flow of bloodthrough a left ventricular assist device, comprising: determining atleast one of ventricular pressure, ventricular volume, and ventricularwall location using the tip of claim 5; and adjusting blood flow througha left ventricular assist device including said cannula, wherein saidblood flow is adjusted to maintain end-diastolic volume or end-diastolicpressure.
 18. The method of claim 17, wherein ventricular volume isdetermined, and wherein ventricular volume is determined using theequation:$V = {\frac{1}{\propto} \times \rho \times L^{2} \times {\left( {G - G^{P}} \right).}}$19. The method of claim 17, wherein ventricular volume is determined,and wherein ventricular volume is determined using the equation:$\frac{1}{1 - {G_{b}/\gamma}}\rho \; {{L^{2}\left( G_{b} \right)}.}$20. The method of claim 17, wherein said blood flow is adjusted tominimize occurrences of negative pressure.